Medical imaging is considered to be one of the most demanding applications of radiographic imaging, at least in terms of resolution and/or contrast. For instance, in the case of mammography it is generally desirable that a given imaging system be capable of obtaining images with a resolution on the order of fifty microns or better. Contrast requirements are also quite demanding since potential lesions and/or suspicious masses may exhibit x-ray attenuation characteristics similar to that of surrounding “healthy” tissue. With early detection of these lesions and/or masses being extremely desirable, enhancement of image resolution and/or contrast continues to be of increasing importance.
Mammography has been performed using both film-based and digital systems. In film-based systems, x-rays signals are generally transmitted through mammary tissue and received at an appropriate screen (e.g., phosphor screen). Light emitted from the screen, due to excitation by the impinging x-rays, is used to expose a light sensitive film. The film is then developed to yield an image of the patient's breast which can be viewed on a light box. By contrast, a full field light-sensitive detector may be utilized in place of the film in digital systems. In this context, “full field” indicates that the field imaged corresponds to the dimensions of the detector, although this may be substantially less than a full image area of interest to a physician, such as a full breast in the case of mammography. This detector outputs an electronic signal that is indicative of the received radiation intensity across a pixel region. In turn, the signal may be connected and processed into a viewable digital image (e.g., on a high-resolution monitor). As may be appreciated, digital image systems are becoming the norm in view of the attendant image storage and processing advantages.
Regardless of the mammography system utilized, granularity and screen noise may tend limit and/or distort the resultant image. For instance, the image contrast has been shown to be significantly affected by scattered radiation. Indeed, as a radiation source transmits a beam towards a tissue, the beam is both attenuated and scattered by the tissue. The scattered radiation that impinges on the detector (e.g. in the case of a digital system) from a path outside a “direct” or substantially straight path from the radiation source to the detector (“primary ray”) is generally undesirable. Accordingly, mammographic images would ideally be generated “scatter-free”.
One type of scatter is Compton scattering, also referred to as incoherent scattering, that typically occurs when an incident x-ray photon ejects an electron from an atom (e.g., a tissular atom), and an x-ray photon of lower energy is “scattered” from the atom. It may be said that relativistic energy and momentum are conserved during Compton scattering and that the scattered x-ray photon generally has less energy and therefore greater wavelength than the original incident photon. By contrast, Rayleigh scattering, also referred to as Thomson or coherent/classical scattering is generally characterized by an x-ray photon interacting with a whole atom so that the photon is scattered with substantially no change in internal energy to the scattering atom, nor to the x-ray photon. In other words, the wavelength of the scattered photon is generally similar to that of the incident photon prior to being scattered.
Various attempts have been made at reducing the incidence of scatter in radiographic imaging. For instance, the use of anti-scatter grids in full field digital imaging as described above has been shown to reduce the amount of scatter displayed in a resultant image. However, use of these grids generally coincides with a need for significant increase in tissue radiation dosages to generate images of desired resolution, signal-to-noise ratio and/or contrast. As another example, various opaque shields have been utilized to enable collection and calculation of an estimated scatter portion of the total beam intensity. However, use of these opaque shields has resulted in portions of the image information being lost (un-exposed, white areas on film). As a way of avoiding these lost areas in the resultant images, one or more subsequent images may be taken without the shields in an additional exposure period(s), and the scatter calculations of the first image can then be applied to the subsequent image(s). However, this obviously requires additional exposure of the patient to radiation doses, and means or process to remove the opaque shields between exposures.